Development of cell sheet constructs for layer by layer tissue engineering using the blood vessel as an experimental model 4

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unpaired t-test) (Figure 5-2). After repeating the grafting process, water contact angle
was found to drop further to 52.0 ± 1.3°, in line with the above observation (p<0.001).
No further improvement could be seen after two rounds of grafting.

Water Contact Angle
Pristine 1x 2x 3x
40
50
60
70
Water Contact Angle

# of Graft layers

Water Contact Angle
degrees
Pristine 70.9±0.3
1 59.6±0.3
2 52.0±1.3
3 51.6±0.4
Figure 5-2: Surface wettability of modified films
Static water contact angle showed a decrease after one round of pAAc grafting, with
another significant decrease with a second round of pAAc grafting. (** denotes
statistically significant difference, p<0.001). No significant change was detected
following subsequent grafting (p>0.05).

5.2.1.3 Quantification of carboxyl density
I used a TBO assay to quantify the amount of carboxyl groups on the surface. PAAc
was grafted onto PCL film surfaces by plasma immobilisation with a 13-fold increase
in surface carboxyl density using Toluidine BlueO (TBO) assay (p<0.01) (Figure 5-3).
This is in contrast to spin-coated polyacrylic acid without plasma cross-linking, which
led to an adsorbed layer that was easily washed off. In line with previous observations,
repeating the spin coating and plasma immobilisation process resulted in increased
grafting yield (p<0.001) to a maximum of 240 nmlol/cm
2
(Figure 5-3).

**
**
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Biaxial stretching is known to influence the morphology and crystal structure of PCL
films, due to fibriallar orientation (Ng 2000; Tiaw 2007). Thus, in a concurrent
experiment, I studied the effects of biaxial drawing on plasma immobilisation yields
of PAAc. A greater grafting yield on biaxially stretched films compared to
unstretched / native films, with 60% higher carboxyl density after two rounds of
grafting (242.4±9.3 vs. 153.7±10.6 nmol/cm
2
, p<0.01, TBO method, Figure 3a).
These results suggest that bi-axial stretching breaks down the lamellae structure,
leading to increased exposure of more polyester linkages for the facilitation of plasma

cross-linking.

Carboxyl Density
Pristine 1x 2x 3x 4x
0
100
200
300
Unstretched
Stretched
Number of Repeat Graftings
Carboxyl Density (nmol/cm
2
)

Figure 5-3: Surface carboxyl density of modified films
Quantification of COOH surface density using Toluidine Blue O Assay demonstrated
a progressive increase in carboxyl density with repeated grafting up to two times, with
stretched films incorporating more carboxyl groups than non-stretched films up to
three repeated treatments. (* denotes statistically significant difference, p<0.01)


*

*

*
*
*
*

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5.2.1.4 Assessment of tensile properties
The choice of the plasma-based method to modify the µXPCL surface was made on
the basis of its advantageous low depth of penetration . However, bioresorbable
materials are by nature highly sensitive to degradation effects. This issue is
compounded in the use of microthin films. Moreover, repeat plasma grafting was used
that may have compromised the properties further. In addition, the modified films
were designed to support conjugation of biomolecules, which necessitated the use of
carbodiimide chemistry. Carbodiimides are commonly used as cross-linkers to
stabilise collagen structures (Nam 2008), and may influence the mechanical behaviour
of the microthin film.

Thus, to evaluate the extent of mechanical damage, I carried out tensile tests to
evaluate the strength of the film following modifications. No significant drop was
observed as a result of plasma treatment, nor was their any trend to suggest
mechanical compromise, irrespective of the number of rounds of PAAc grafting. In
light of the results thus far, subsequent experimentation was conducted on µXPCL
grafted with two layers of PAAc (Henceforth referred to as PCL-PAAc).

Surprisingly, a drop in ultimate tensile strength, albeit insignificant, was observed for
carbodiimide treated films. It suggests that some form of degradation may have taken
place, and will warrant further studies in the future. No similar trend could be
observed in the yield strength (Figure 5-4).

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UTS
Pristine 1x 2x 3x 2x-EDC
0
10
20
30
40
50
60
70
80
90
100
110
120
Stress, MPa
Yield Strength
Pristine 1x 2x 3x 2x-EDC
0
5
10
15
20
25
30

35
40
45
Yield Strength, MPa

Sample UTS σ
y
Pristine 90.0 ± 7.4

32.3 ± 5.5

PCL-1x 92.4 ± 11.9

37.3 ± 5.5

PCL-2x 85.2 ± 9.8

29.4 ± 6.0

PCL-3x 92.2 ± 10.6

38.5 ± 5.2

Activated PCL-2x 69.0 ± 9.9

25.5 ± 7.0


Figure 5-4: Mechanical properties of modified films
Tensile testing of µXPCL-PAAc films demonstrate no trend nor significant difference

in (a) ultimate tensile strength and (b) yield strength as a result of plasma modification
(p>0.05). A drop in tensile properties occurred as a result of the carbodiimide
activation process, but the decrease was not statistically significant.

5.2.1.5 Plasma immoblisation of Polyethyleneimine (PEI)
It is possible to confer greater functionality to the film surface by varying the nature
of the grafted polymer. This includes polyols for the introduction of chemically useful
hydroxyl groups (Bures 2001) or even polymers with specific configurations, such as
star-shaped dendrimers (Won 2003). Here, I studied the immobilisation of PEI to
provide amine groups.

PEI was successfully grafted using a similar process. However, because of the lack of
reliability of amine quantification techniques, I analysed the PEI modified surface
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with XPS. Unlike PCL and PAAc, PEI contains nitrogen, and will show up on the
XPS scan through the introduction of the N1s peak (Figure 5-5).
0200400600800100012001400
0
0.5
1
1.5
2
2.5
3
3.5

4
4.5
x 10
4
PCL017.spe
Binding Energy (eV)
c/s
-O KLL
-C KLL
Atomic %
C1s 72.2
O1s 22.6
N1s 4.6
S2p 0.6
-O1s
-N1s
-C1s
-S2s
-S2p

Figure 5-5: XPS survey of PEI-immobilized film
XPS wide survey spectra of PEI grafted µXPCL films shows the introduction of an
N1s peak due to the presence of amines.

5.3 Conjugation of Biomolecules
In this section, I investigated the possibility of conjugating bioactive molecules onto
the modified films for possible tissue engineered blood vessel applications.
Conjugation of a wide variety of biomolecules to blood contacting surfaces has been
studied in vascular tissue engineering, which can be broadly classified under three
categories: adhesion proteins, anticoagulants and antibodies. I had chosen a candidate

from each category, and demonstrated the conjugation onto modified µXPCL surfaces.

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5.3.1 Conjugation Of Heparin
The conjugation of heparin onto blood contacting surfaces has been well-studied.
Beyond anti-coagulatory properties, heparin has been shown to possess growth factor
binding sites and has been shown to improve the efficacy of growth factors (Mitsi
2008). It has also been shown to reduce complement activation (Girardi 2005).
Moreover, heparin-like proteoglycans are found abundantly on basement membranes
and on the surface of EC, suggesting that heparin will be suitable for lining blood
contacting surfaces.

In my initial experiments with grafting heparin onto PCL-PAAc, low and inconsistent
engraftment yields were observed. This may have arisen due to inefficiencies in the
deprotonation of the heparin sodium salt. Consequently, I opted to graft heparin onto
the PEI modified PCL instead. XPS scans reveal the introduction of an S2p peak due
to the incorporation of the highly sulphated heparin. Presence of the N1s peak was
due to the PEI.
0200400600800100012001400
0
0.5
1
1.5
2
2.5

3
3.5
4
4.5
x 10
4
PCL014.spe
Binding Energy (eV)
c/s
-C KLL
-O KLL
Atomic %
C1s 66.7
O1s 23.0
N1s 8.6
S2p 1.7
-O1s
-N1s
-C1s
-S2s
-S2p

Figure 5-6: XPS survey of heparin conjugated PCL-PEI film
XPS wide survey spectra of modified film following conjugation of heparin,
demonstrating the increased sulphur content.
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5.3.2 Conjugation Of Collagen
Collagen is commonly used for surface modification in tissue engineering
applications to improve cellular adhesion. Collagen contains RGD (Arg-Gly-Asp)
sites that are recognised by cells via integrins. I proceeded to conjugate collagen onto
PCL-PAAc films by carbodiimide chemistry. XPS scans demonstrate that collagen
was successfully grafted onto the film surface, as evidenced by the N1s peak.
However, it was also noted that significant contaminants were present. The source of
collagen used varies greatly from batch to batch due to its direct isolation from animal
products (oxen tails), and may have contributed to the contamination observed here.

0200400600800100012001400
0
0.5
1
1.5
2
2.5
3
3.5
4
4.5
5
x 10
4
KS05.spe
Binding Energy (eV)
c/s
-Na KLL
Atomic %

C1s 68.5
O1s 20.1
N1s 8.9
Na1s 2.5
-Na1s
-O1s
-N1s
-C1s

Figure 5-7: XPS survey of collagen grafted PCL-PEI film
XPS wide survey spectra of PEI grafted µXPCL films shows the introduction of an
N1s peak due to the presence of amines. Some contamination was observed, possibly
due to purity issues at source.

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5.3.3 Conjugation Of Fluorescent Antibodies
To visualise the distribution of the conjugate, I proceeded to graft fluorescently
labelled biomolecules on the surface. Commercial fluorescently labelled antibodies
were used because they are consistent, well-characterised and optimised to reduce
signal-to-noise ratio. The antibodies were grafted on successfully using carbodiimide
chemistry. In contrast, without carbodiimide conjugation, nonspecific adsorption
occurred, resulting in the diffuse staining. Following washing for 72 hours under
representative physiological conditions (mechanical shaking in physiological saline at
body temperature), most of the antibodies on the untreated film were washed off,
leaving some residual adsorbed antibodies. Labelled activated films continued to

express fluorescence.

Figure 5-8: Micrographs of fluorescent-label conjugated PCL-PAAc
Epifluorescent micrographs of films conjugated with Alexafluo 488 goat anti-mouse
secondary antibody. The label was conjugated onto activated films, and fluorescence

5.3.4 Conjugation Of Monoclonal CD34 Antibody
Antibodies have previously been proposed for use as “capture surfaces” in blood
contacting applications to recruit albumin from the circulation, thus forming a self-
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renewing passivated surface. More recently, antibodies against specific cell types
have been proposed. In particular, surfaces immobilised with antibodies against CD34
have been studied for recruitment of circulating endothelial progenitor cells. This
represents a viable option for vascular graft design.

5.3.4.1 XPS analysis of PCL-CD34
CD34 antibodies were grafted onto PCL-PAAc surfaces (PCL-CD34), and verified by
XPS (Figure 5-9). Wide survey XPS scans show increased O:C ratio following PAAc
grafting, and then the introduction of an N1speak following the conjugation of the
CD34 antibody, indicating presence of the protein. High resolution surveys of the C1s
core spectra revealed the presence of peptide bonds. I also noted an increase in the
percentage of carboxyl containing groups on PCL-PAAc, as compared to µXPCL.
This dropped following the conjugation of CD34 antibody, as the carboxyl groups
were consumed in the conjugation process.
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Figure 5-9 XPS analysis of PCL-CD34
Engraftment of CD34 antibody a) XPS wide survey spectra of the modified films
demonstrating successful engraftment (b) Relative intensity of the deconvoluted C1S
spectra of µXPCL films shows increase in carboxyl groups following PAAc
engraftment,followed by introduction of peptide groups following CD34 antibody
conjugation

Atomic Ratio (%)
-C-H -C-O -O-C=O -C-N -O-C-N
Binding
Energy (eV)
~284.6 ~286.2 ~288.6 ~285.8 ~287.4
Pristine 71.2 15.8 13.0 - -
µXPCL-PAAc 70.7 11.8 17.5 - -
µXPCL-CD34 63.7 10.4 11.5 10.3 4.1
282284286288290
284286288290
284286288290
µXPCL
PCL-PAAc
PCL-CD34
(b)
XPS

0250500750100012501500
µXPCL
µXPCL-PAAc
µXPCL-CD34
µXPCL-PAAc-NHS/EDC
Wavelength
(a)
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5.3.4.2 AFM imaging
AFM imaging shows fibrillar morphology typical of stretched spherulites on PCL
films as a result of biaxial stretching (Figure 5-10 a, b). On PCL-PAAc, the fibrils
were partially obscured by the engrafted layer, resulting in a drop in roughness
(3.2±1.1 nm vs 6.7±1.4 nm, p<0.001) (Figure 5-10c). PCL-CD34 showed a distinct
topography in Figure 5-10d where the conjugated CD 34 antibodies were found to
adopt a globular morphology. Roughness measurements show that PCL-CD34 is
smoother than PCL (4.5±0.6 nm, p<0.01), and similar to PCL-PAAc where the
difference was insignificant (p>0.05).
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PCL PCL-PAAc PCL-CD34
Roughness 6.74±1.43 nm 3.24 ± 1.09 nm 4.46 ± 0.58 nm

Figure 5-10: AFM analysis of modified film surfaces
(a) Three dimensionally rendered images (Scan area 5 µm x 5 µm) of the film surfaces
following surfaces. PCL film surfaces demonstrate aligned fibres arising from the
biaxial stretching process, which are covered by the grafted polyacrylic acid layer and
the conjugated CD34 antibody (b) Similarly, images of the film surfaces captured
using amplitude signal (Scan area 1 µm x 1 µm) shows that the surface is covered by
PAAc. Conjugated CD34 antibodies show up as globular structures. Surface
roughness was evaluated and presented in (c).

5.3.4.3 Carboxyl density
Following engraftment of PAAc to PCL, carboxyl density increased from 2.6±1.0
nmol/cm
2
to 170.8±4.4 nmol/cm
2
on TBO assay. The addition of CD34 antibody
conjugation to raised the carboxyl density further to 216.7±7.9 nmol/cm
2
(p<0.001).

(b)
(a)
PCL

PCL
-

PAAc

PCL
-
CD34

PCL

PCL
-
PAAc

PCL
-
CD34

(c)
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TBO Assay
0
50
100
150
200
250

***
***
PCL PCL-PAAc
PCL-CD34
COOH Concentration (nM/cm
2
)



Figure 5-11: Surface carboxyl density of surface modified films
Surface carboxyl density of pristine and modified PCL films as assessed by modified
Toluidine BlueO assay. *** denotes significantly increased carboxyl density over
pristine PCL films (p<0.001)

5.3.4.4 Stability
To visualise the distribution of the antibodies, I imaged the films using
immunohistochemical techniques. The immobilised CD34 antibodies picked up the
label, staining the films green. In contrast, PCL-PAAc films showed little staining
after the antibodies were washed off. Subjecting PCL-CD34 to washing under
physiological conditions did not deplete the antibodies on the surface, as evidenced by
a retention of fluorescence intensity following 72 hours.

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Figure 5-12: Fluorescent imaging of PCL-CD34 films
(a) When films were immunostained with fluorescent labels, control PCL-PAAc film
did not exhibit fluorescence, as compared to (b) µXPCL-CD34, showing uniform
immobilisation of bound antibodies on the film (c) Films were washed under
mechanical shaking, and measurements of samples retrieved after 24 and 72 hours
Residual Mean Fluoresence Intensity on the films was retained, showing antibodies
remain stably bound. NS denotes no significant difference (p>0.05)

Mean Fluoresence Intensity
24h 72h Ctrl
0
10
20
30
40
50
(c)
PCL-CD34
(a)

(b)

NS
PCL-PAAc
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5.4 Discussion
5.4.1 Summary Of Results
µXPCL surfaces are inherently hydrophobic and bioinert, and surface modification is
necessary to improve the biocompatibility of the material. Wet chemical methods are
inappropriate, due to the penetrative nature of such treatments and the microthin
nature of the films. In particular, bioresorbable polyesters, such as PCL, are highly
susceptible to chemical and thermal degradation, and consequently, mechanical
properties are often compromised as a result of the modification process. Furthermore,
large amounts of chemicals are typically involved, which are less environmentally
friendly, and potentially toxic if not removed properly (Abidi 2004). Thus, I applied
plasma-based techniques to immobilise hydrogel layers on the µXPCL film surfaces,
and demonstrated that the mechanical properties were not significantly affected as a
result of surface modification. In particular, polyacrylic acid could be immobilised on
the surface, thus achieving functionalisation of the surface with carboxyl groups, with
corresponding reduction in static water contact angle. The significant loading of
functional groups indicates that the modified surfaces could be used for subsequent
conjugation of biomolecules. Furthermore, the process could be adjusted to get a
range of carboxyl densities, and corresponding range of hydrophilicity for specific
applications.

I then proceeded to demonstrate the conjugation of biomolecules onto the hydrogel-
modified surfaces. Heparin, collagen and antibodies were all successfully conjugated.
In particular, I have successfully conjugated CD34, a endothelial progenitor cell
selective-antibody, to the films. The conjugated biomolecules remain attached even
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174

after thorough repeated washing. My results confirm the feasibility of the use of
plasma based methods for the surface modification of µXPCL films without affecting
its tensile strength.

5.4.2 Critical Assessment
5.4.2.1 Plasma immobilisation of hydrogels
Similar plasma techniques have previously been used to deposit PAAc layers onto
bioresorbable nanofiber mats (Park 2007). However, such deposited layers require
mechanical interlocking with asperities on the substrate surface, and consequently, are
unstable when deposited on smooth substrates (Kato 2003). Although the plasma
immobilisation method described here has previously been shown to be inefficient
due to etching effects (Terlingen 1993; Terlingen 1994). Nitschke et al achieved
success with the plasma immobilisation of various hydrogels onto both PET and PE
surfaces, suggesting a dependence of grafting efficiency on the following parameters:
(1) wettability of substrate (2) thickness of hydrogel layer and (3) molecular weight of
PAAc.

Using the Nitschke method, I found that PAAc could similarly be engrafted onto the
µXPCL surface, and that it was possible to increase the grafting yield simply by
repeating the process. The plasma immobilisation occurs via radical mechanism (Lens

1998) which first involves the generation of free radicals via chain scission and
hydrogen abstraction, followed by crosslinking of the adsorbed polymeric layer. This
allows the subsequent immobilisation of additional PAAc layers by attachment to
underlying layers. The process, however, is limited by etching (Terlingen 1994),
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resulting in maximum carboxyl density achieved after three repeats of the process.
Using this method, I achieved a range of carboxyl densities from 100 to 200 nmol/cm
2
.
This figure compares favourably against other groups using other methods, and shown
to be suitable for the subsequent conjugation of biomolecules (Sano 1993; Grondahl
2005; Du 2006).

5.4.2.2 Conjugation of biomolecules
In my experiments with heparin and collagen-conjugated films, I often found the
presence of contaminants. A likely cause for the contaminations is the source.
Currently, many biomolecules are derived from animal tissue, and despite stringent
quality control, valid concerns have been raised on the quality, purity and
predictability of such products. Recombinant proteins are increasingly being studied,
providing a possible solution to these concerns in the future (Ito 2006). Thus, I have
concentrated my investigations on the purer and better characterised monoclonal
antibodies, CD34 in my case, synthesised through hybridomas, which should have
more predictable results.

To promote the adhesion of endothelial progenitor cells (EPC) on µXPCL-PAAc, in

order to generate a confluent endothelium in vivo, I demonstrated the use of EPC-
selective CD34 antibodies immobilised on the substrate surface to anchor EC types
(Shown below in 6.2.2.2). Endothelial cell types are known to secrete extracellular
matrix (ECM) proteins slowly (Divya 2007), and this anchorage step may be crucial
in stabilising the cell during early stages of cell adhesion. CD34 antibodies offer the
additional advantage of cell selection. EPC are highly proliferative and shown to
integrate into host vasculature, traits that are favourable for regenerating the
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endothelium, but are lost in terminally differentiated cells. Capture of CD34 positive
cells may thus improve the healing and tissue integration process. Although
ubiquitous extracellular matrix proteins such as fibronectin or vitronectin are more
commonly used in tissue engineering, such molecules are known to lead to
undesirable biological responses such as thrombosis (Stephan 2006).

In this part of the study, however, I have not studied the activity of the immobilised
molecule, nor did I manage to quantify the surface density of CD34 antibodies.
Carbodiimide induced cross-linking has been associated with loss of protein activity,
due to the non-specific nature of the process. This leads to excessive cross-linking,
resulting in changes in protein conformation and loss of activity (Camarero 2008).
Furthermore, the cross-links may occur at or near the active site, obscuring the ability
of the antibody to identify the cell antigen. In my experiments, I had used titration
methods to assess the minimum amount of carbodiimide required to facilitate
conjugation, in order to reduce the frequency of cross-linkages. However, there are no
assays available to study the biological activity of the surface, and I proceeded to
characterise the biological response to the modified surface instead. The use of

recombinant affinity tags and site specific covalent immobilisation have recently
emerged as sophisticated methods to address this issue (Holt 2000; Olsson 2000), and
may be worth pursuing in the future. Another key limitation in the use of
carbodiimides and other zero-length cross linkers is steric hindrance. This reduces
mobility of the bound protein, and may disrupt biological interactions. If necessary,
the use of linkers and spacers, such as polyethylene glycol, may be used in the future
(Kuhl 1996).
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177












6 Surface Modification to Improve Biocompatibility
Chapter 6. Surface Modification to Improve Biocompatibility


178

6.1 Introduction

Common host responses that tissue engineers try to modulate include inflammation,
healing and cell adhesion. I have described in Chapter 3 the development of microthin,
bioresorbable films with slow degradation kinetics to reduce inflammatory responses.
Cell seeding is commonly employed to accelerate the healing process, and I have
detailed in Chapter 4 my studies on candidate vascular progenitor cells. I had also
described in Chapter 5 modification methods to engineer the film surfaces which can
be conferred with biological functionality through conjugation of biomolecules. These
developments contribute towards improved biocompatibility of the tissue engineering
scaffolds.

In the specific context of vascular tissue, two other important classes of tissue
responses remain. First, possible issues of thrombogenesis must be addressed. Acute
failures of vascular graft occurring within the first three months of implantation are
caused by technical complications, primarily thrombosis. The graft is in constant
contact with blood, and must perform the homeostatic regulation of blood clot
formation. Next, vascular grafts typically fail in the mid-term because of re-occlusion
arising from neointimal hyperplasia. Existing vascular grafts are stiff, resulting in the
mechanical mismatch that is believed to direct smooth muscle cells from the media
towards a hyperproliferative synthetic phenotype.

I have described in Chapter 1 layer-by-layer approach that mimic vascular
architecture. I describe in this chapter the process of film modification and
optimisation for improved biocompatibility. Because vascular tissue is organised into
stratified tunicae, I had suggested the customisation of surfaces to support cell types
Chapter 6. Surface Modification to Improve Biocompatibility


179

specific to the tunicae. Thus, I studied the adhesion of perivascular cells to modified

PCL film surfaces for the generation of the perivascular compartment. I then
proceeded to engineer a surface for the reconstitution of the tunica intima. The intima
is constantly exposed to fluid shear stresses, and consequently, suitable methods to
anchor the cultured EC are necessary. Additionally, the surface may be denuded as a
result of the surgical procedure or natural sloughing in vivo, thus the underlying
material must be designed to be blood compatible. These factor led to my selection of
CD34 antibodies as a suitable conjugand, and I detail here my studies on cell adhesion,
as well as blood compatibility.

Finally, I studied the assembly of cells and film to form a layered structure. The
tunicae in vascular tissue are unique in the composition of cells: each compartment is
only populated by the resident cell type. Thus, endothelial and smooth muscle cells
reside as distinct populations in the tunica intima and tunica media respectively,
separated by a basement membrane. It has been postulated that disruptions of this
configuration contribute towards changes in cell phenotype; in particular, disrupted
endothelium contributes towards transformation of smooth muscle cells in the media.
Consequently, it is imperative that the different cell types are maintained distinct from
each other prior to the formation of confluent layers. I describe here my co-culture
studies towards the generation of distinct layers, reminiscent of the native vessel.

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6.2 Biological Responses To Surface Modified
µ
XPCL Films
6.2.1 Blood Compatibility Experiments
I chose three aspects of blood clotting to evaluate: Contact activation, platelet

activation, as well as whole blood responses.

6.2.1.1 Contact activation
TEG profiles generated for each group was found to have classical cigar shapes indicating
functional clotting typical of normal haemostasis
(Figure 6-1).
There were no significant
differences in A° and MA across the groups (p>0.05). The TEG profile for PCL was found to
be closest to that of the negative control, with comparable r and MA values. The k time,
however, was found to be significantly reduced (12.7 min vs 7.5 min, p<0.05). Following
PAAc engraftment, contact activation was significantly increased compared to PCL, with
significantly reduced r (11.6 min, p<0.001) and k times (3.5 min, p<0.05), but no significant
difference in MA. There was a trend towards an increased A° in PCL-PAAc over PCL films,
albeit insignificantly. Subsequent conjugation of CD34 antibody increased both r and k times
(19.3 min and 6.9 min respectively, p<0.001) and reduced A° to levels comparable with PCL.
Chapter 6. Surface Modification to Improve Biocompatibility


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Figure 6-1: TEG profiles of blood in contact with modified film surfaces.
Glass was included as a positive control and citrated blood was used as a negative
control. (a) Profiles demonstrate similar cigar shape, indicating functional
haemostasis. Tracings taken from left to right: Glass, PCL-PAAc, PCL, PCL-CD34
and Negative Control.


Figure 6-2: Graphs of TEG parameters

Comparison of (a) Reaction time, R (b) Clotting time, k and (c) Maximum Amplitude,
MA. Data are means ± SD, n=4. (*): p<0.05, (NS): Not significant

TEG - Angle
Neg Ctrl PCL PCL CD34 PCL PAAc Glass
0
20
40
60
80
Degrees
TEG - MA
Neg Ctrl PCL PCL CD34 PCL PAAc Glass
0
20
40
60
80
mm
TEG - K
Neg Ctrl PCL PCL CD34 PCL PAAc Glass
0
5
10
15
20
***
*
NS
NS

Min
TEG - R
Neg Ctrl PCL PCL CD34 PCL PAAc Glass
0
10
20
30
*
NS
NS
Min
Chapter 6. Surface Modification to Improve Biocompatibility


182

6.2.1.2 Platelet Adhesion
Widespread and uniform platelet attachment could be found on pristine PCL surface,
with 2.5 times as many platelets attaching to PCL than glass (p<0.001) (5,650 vs
2,167, Figure 6-3). PAAc engraftment resulted in a 9-fold decrease in number of
adhered platelets compared to PCL (620, p<0.001). The addition of CD34 antibodies
resulted in a doubling of attached platelets, although this difference was statistically
insignificant (1,389, p>0.05) (Figure 6-3).

Platelet morphology was used as an indication of extent of activation in response to
the samples surfaces (Figure 6-4). Adhered platelets found on the pristine PCL films
had developed pseudopodia-like structures from the platelet body, with spread
hyaloplasm and well-established fibrin networks branching out to the surrounding
area, (yellow arrows, Figure 6-4b), similar to those observed on glass (Figure 6-4a).
In contrast, platelets adhering to PCL-PAAc (Figure 6-4c) appeared to be dendritic

without evidence of flattening. Similarly on PCL-CD34 (Figure 6-4d), platelets were
largely dendritic, with some intermediate pseudopodia.
Chapter 6. Surface Modification to Improve Biocompatibility


183

Platelet Adhesion
G
las
s
PC
L
P
AA
c
CD
34
0
2000
4000
6000
8000
NS
***
Platelet count (per mm
2
)

Figure 6-3: Platelet adhesion studies

Scanning Electron Microscope images were obtained of surfaces after exposure to
platelet rich plasma (500x magnification) Platelet counts demonstrated significantly
higher number of platelets adhering to PCL surface as compared to PCL-PAAc and
PCL-CD34 (*** denotes p<0.001, NS: p>0.05)


Glass

PCL
PCL-PAAc PCL-CD34